U.S. patent application number 13/810648 was filed with the patent office on 2013-08-15 for ultrasonic horn actuated microprobes based self-calibrating viscosity sensor.
This patent application is currently assigned to CORNELL UNIVERSITY. The applicant listed for this patent is Ramkumar Abhishek, Amit Lal. Invention is credited to Ramkumar Abhishek, Amit Lal.
Application Number | 20130205875 13/810648 |
Document ID | / |
Family ID | 45470079 |
Filed Date | 2013-08-15 |
United States Patent
Application |
20130205875 |
Kind Code |
A1 |
Lal; Amit ; et al. |
August 15, 2013 |
ULTRASONIC HORN ACTUATED MICROPROBES BASED SELF-CALIBRATING
VISCOSITY SENSOR
Abstract
An ultrasonic or acoustic viscosity sensor or viscometer is
provided that can be used to accurately measure viscosity for fluid
samples of less than 1 .mu.l in volume. Methods for measuring
viscosity for fluid samples of less than 1 .mu.l in volume are also
provided. The viscosity sensor and methods based thereon enable
simultaneous measurement of bulk and dynamic (shear-rate dependent)
viscosity of a non-Newtonian fluid. Bulk and dynamic viscosity of
the non-Newtonian fluid can be measured simultaneously without
separating constituents of the fluid, and thus distinguishing the
effect of constituents on the viscosity. Dynamic viscosity of the
non-Newtonian fluid can be estimated at varying shear rates, to
study the deformability of the constituents of the fluid as a
function of shear rate.
Inventors: |
Lal; Amit; (Ithaca, NY)
; Abhishek; Ramkumar; (Mountain View, CA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Lal; Amit
Abhishek; Ramkumar |
Ithaca
Mountain View |
NY
CA |
US
US |
|
|
Assignee: |
CORNELL UNIVERSITY
Ithaca
NY
|
Family ID: |
45470079 |
Appl. No.: |
13/810648 |
Filed: |
July 14, 2011 |
PCT Filed: |
July 14, 2011 |
PCT NO: |
PCT/US11/44032 |
371 Date: |
April 30, 2013 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
61365159 |
Jul 16, 2010 |
|
|
|
Current U.S.
Class: |
73/54.41 |
Current CPC
Class: |
G01N 2291/02818
20130101; G01N 29/2437 20130101; G01N 11/16 20130101; G01N 29/036
20130101 |
Class at
Publication: |
73/54.41 |
International
Class: |
G01N 11/16 20060101
G01N011/16 |
Goverment Interests
STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT
[0002] The disclosed invention was made with government support
under Contract No. 5R01HL073644-0, entitled "MEMS sensors for
arrhythmia detection and intervention" from the National Institutes
of Health. The government has rights in this invention.
Claims
1. An ultrasonic or acoustic viscosity sensor for measuring fluid
viscosity or fluid flow of a fluid in real time comprising: an
ultrasonic or acoustic actuator 10; at least one microprobe 20 for
sensing ultrasound induced motion; and means for generating motion
via coupling between the actuator and the at least one
microprobe.
2. The ultrasonic or acoustic viscosity sensor of claim 1
comprising a substrate 25.
3. The ultrasonic or acoustic viscosity sensor of claim 2 wherein
the actuator couples motion to the at least one microprobe through
the substrate.
4. The ultrasonic or acoustic viscosity sensor of claim 1 wherein
the actuator is driven at at least one of its mechanical resonance
frequencies.
5. The ultrasonic or acoustic viscosity sensor of claim 4 wherein
the at least one microprobe is driven at at least one of its
mechanical resonance frequencies by the actuator.
6. The ultrasonic or acoustic viscosity sensor of claim 1 wherein
the actuator is a silicon ultrasonic longitudinal mode
actuator.
7. The ultrasonic or acoustic viscosity sensor of claim 6 wherein
the silicon ultrasonic longitudinal mode actuator has varying
cross-section for amplifying motion imparted to the at least one
microprobe.
8. The ultrasonic or acoustic viscosity sensor of claim 1 wherein
the actuator is actuated by a piezoelectric element 16.
9. The ultrasonic or acoustic viscosity sensor of claim 8 wherein
the piezoelectric element is adhesively attached.
10. The ultrasonic or acoustic viscosity sensor of claim 8 wherein
the piezoelectric element is a thin piezoelectric film.
11. The ultrasonic or acoustic viscosity sensor of claim 1 that
operates in acoustic frequencies between 10 Hz to 20 kHz.
12. The ultrasonic or acoustic viscosity sensor of claim 1 that
operates in mid ultrasonic frequencies between 20 kHz to 500
kHz.
13. The ultrasonic or acoustic viscosity sensor of claim 1 that
operates in upper ultrasonic frequencies between 500 kHz to 10
MHz.
14. The ultrasonic or acoustic viscosity sensor of claim 1
comprising means for measuring damping of the mechanical vibrations
of the at least one microprobe 200 in the fluid of interest.
15. The ultrasonic or acoustic viscosity sensor of claim 14 wherein
the means for measuring damping of the mechanical vibrations
comprises at least two strain gauges or piezoresistors located at
the junction of the silicon horn and the microprobes, wherein at
least one of the strain gauges or piezoresistors measures the
amplitude of motion of the at least one microprobe and wherein at
least one of the strain gauges or piezoresistors measures the
motion of the actuator.
16. The ultrasonic or acoustic viscosity sensor of claim 15 wherein
at least one of the strain gauges or piezoresistors is a
polysilicon resistor.
17. The ultrasonic or acoustic viscosity sensor of claim 15 wherein
at least one of the strain gauges or piezoresistors is integrated
or positioned at the junction of the silicon horn and the
microprobe.
18. The ultrasonic or acoustic viscosity sensor of claim 15 wherein
at least one of the strain gauges or piezoresistors has a nominal
resistance of about 10 k.OMEGA..
19. The ultrasonic or acoustic viscosity sensor of claim 1
comprising means for measuring immersion depth of the microprobe
230 in the fluid.
20. The ultrasonic or acoustic viscosity sensor of claim 19 wherein
the means for measuring immersion depth is based on measurement of
capacitance between two electrodes.
21. The ultrasonic or acoustic viscosity sensor of claim 20 wherein
the means for measuring immersion depth is a capacitance-based
immersion depth sensor.
22. The ultrasonic or acoustic viscosity sensor of claim 21 wherein
the capacitance-based immersion depth sensor comprises means for
distance-coding capacitance.
23. The ultrasonic or acoustic viscosity sensor of claim 22 wherein
the means for distance-coding capacitance is formed by at least two
metal traces on the capacitance-based immersion depth sensor.
24. The ultrasonic or acoustic viscosity sensor of claim 1
comprising computational circuitry functionally connected to the
microprobe.
25. The ultrasonic or acoustic viscosity sensor of claim 1
integrated into a strip format.
26. The ultrasonic or acoustic viscosity sensor of claim 25 wherein
the strip format is a blood-glucose measurement strip.
27. The ultrasonic or acoustic viscosity sensor of claim 1
comprising: a silicon horn 15; a piezoelectric actuator element 16;
and a multi-sensor microprobe 20 for sensing ultrasound induced
motion, wherein: the multi-sensor multiprobe comprises: at least
one polysilicon strain gauge or piezoresistor 200 connected in a
Wheatstone bridge 250 configuration, and a capacitance-based
immersion depth sensor 230.
28. The ultrasonic or acoustic viscosity sensor of claim 1 that is
mid-frequency resonant and wherein the mid-frequency is 20-500
kHz.
29. The ultrasonic or acoustic viscosity sensor of claim 1 that is
high-frequency resonant and wherein the high-frequency is 500
kHz-10 MHz.
30. The ultrasonic or acoustic viscosity sensor of claim 1
comprising at least one multi-sensor microprobe 20.
31. The ultrasonic or acoustic viscosity sensor of claim 30 wherein
the at least one multi-sensor microprobe comprises a plurality of
strain gauges or piezoresistors 200.
32. The ultrasonic or acoustic viscosity sensor of claim 1 wherein
the microprobe comprises a plurality of capacitance-based immersion
depth sensors 230.
33. The ultrasonic or acoustic viscosity sensor of claim 30 wherein
the at least one multi-sensor microprobe 20 comprises a plurality
of capacitance-based immersion depth sensors 230.
34. A method for measuring fluid viscosity or fluid flow of a fluid
of interest comprising the steps of: providing the ultrasonic or
acoustic viscosity sensor 1 of claim 1; and measuring viscous
damping of non-linear flexural vibrations of the microprobe 20
induced by the .lamda./2 longitudinal resonance of the silicon horn
15.
35. The method of claim 34 wherein the non-linear flexural
vibrations of the microprobe comprise longitudinal and flexural
oscillating motions.
36. The method of claim 34 wherein the step of measuring viscous
damping comprises measuring viscous damping at multiple
frequencies.
37. The method of claim 36 wherein a first of the multiple
frequencies is the actuator 10 longitudinal mode frequency and a
second of the multiple frequencies is the microprobe 20 resonance
frequency.
38. The method of claim 34 comprising the step of calculating the
frequency and magnitude of a signal corresponding to the
multi-frequency excitation from the Fast Fourier Transform (FFT) of
the amplified Wheatstone bridge 250 voltage signal corresponding to
the motion of a cantilever of the at least one microprobe 20.
39. The method of claim 34 comprising the step of calibrating
immersion depth.
40. The method of claim 39 wherein the calibrating step comprises
measuring capacitance at discrete steps.
41. The method of claim 34 comprising the step of determining the
shear decay length (.delta.) at the actuator resonance, wherein
.delta. is given by .delta. = ( 2 .eta. .omega..rho. ) 1 2
##EQU00003## where .eta. and .rho. are the fluid viscosity and
density, respectively, and .omega.=2.pi.f is the angular frequency
of flexural vibration of the at least one microprobe.
42. The method of claim 34 comprising the step of measuring the
viscosity of the fluid of interest at different shear rates.
43. The method of claim 42 comprising varying the piezoelectric
transducer (PZT) actuation voltage.
44. The method of claim 34 comprising the step of assessing
coagulation time of the fluid of interest.
45. The method of claim 34 comprising the step of measuring the
coagulation of the fluid of interest.
46. (canceled)
Description
CROSS-REFERENCE TO RELATED APPLICATIONS
[0001] This application claims priority to and the benefit of
co-pending U.S. provisional patent application Ser. No. 61/365,159,
entitled "Ultrasonic horn actuated microprobes based
self-calibrating viscosity sensor," filed Jul. 16, 2010 which is
incorporated herein by reference in its entirety.
1. TECHNICAL FIELD
[0003] The present invention relates to fluid viscosity and flow
sensors. The invention also relates to devices for measuring
viscosity and flow of blood.
2. BACKGROUND OF THE INVENTION
[0004] Medical diagnosis of disease often requires determining the
viscosity of body fluids, very importantly that of blood. Blood
rheology in vivo can be used for prediction of blood flow through
capillaries of varying sizes, while its in vitro rheology can be
used for diagnosis and monitoring of diseases. Also, changes in
blood rheology have been seen to contribute to or aggravate
cardiovascular disorders such as myocardial infarction and
hypertension. Clinical hemorheological tests measure the red blood
cell (RBC or erythrocyte) aggregation by the erythrocyte
sedimentation rate (ESR) and the plasma viscosity. Physiological
conditions leading to alteration of plasma proteins (such as
fibrinogen) results in a change in the ESR and plasma viscosity.
Correlation to blood rheology can be used as a good indicator for
prognosis, diagnosis and/or monitoring disease processes such as
cancer, cardiovascular disease, etc. For example, inflammatory
biomarkers (fibrinogen) and plasma viscosity have been observed to
be significantly associated with death due to myocardial infarction
or coronary heart disease. Incidence of type 2-diabetes has been
shown to be strongly related to high whole blood viscosity (WBV),
thus allowing a measure of blood viscosity to be a prognosis tool.
Also, monitoring plasma viscosity can help in the prediction and
diagnosis of plasma hyperviscosity syndrome and sickle cell
disease. Additionally, a measure of blood viscosity can help in
monitoring interventions to alter blood viscosity such as diet,
alcohol, pharmaceuticals etc.
[0005] Viscosity meters use various fluid-mechanical instruments to
move the fluid applying a shear stress layer which generates a
viscosity modified damping or flow, which are measured
electronically, optically, etc. One device known in the art that
can be used to measure viscosity is a mechanical resonator loaded
with a liquid. The implied velocity field into the fluid from the
resonator results in bulk and shear motion induced damping, both
affecting the quality factor of the resonance, the resonance
frequency, and the resonant motion amplitude, among other
variables. One aspect of medical relevancy of blood viscosity is
the fluidic resistance upon parts moving at displacement greater
than several red blood-cell diameters. A viscosity sensor for blood
should be able to move the blood cells individually such that the
interaction of the blood cells with the plasma is sampled
adequately. This greatly limits the use of high frequency MEMS
resonators where displacements are in the deep sub-micron regime,
for estimating viscosity by monitoring the viscous damping. On the
opposite extreme are low frequency resonant devices that can become
bulky owing to the low spring constant k or high mass m required to
make the resonant frequency low. At low frequencies the size of the
resonators approaches that of commercial tabletop viscometers
requiring tens of milliliters of blood.
[0006] Furthermore, with micro or millimeter scale resonators, the
total contact area of the sample with the probes strongly
determines the loss of mechanical energy into the fluid from the
actuator. Hence, it is important to measure the contact area to
measure viscosity accuracy. In typical sensors, the sensor area is
typically much smaller than the fluid volume completely submerging
the sensor in the fluid. However, in cases where tiny amounts of
samples are present, this condition is hard to maintain.
[0007] There is therefore a need in the art for a fluid viscosity
and flow sensor for use as a prognosis, diagnosis and/or monitoring
device. There is also a need for an instrument to measure body
fluid viscosity at varying shear rates in real-time (for example,
before a blood sample coagulates) using <1 .mu.l fluid samples.
Such an instrument, together with other biosensor(s) such as
glucose sensors for diabetic patients, could serve as an invaluable
tool for rapid monitoring of an individual's physiological
function.
[0008] Citation or identification of any reference in Section 2, or
in any other section of this application, shall not be considered
an admission that such reference is available as prior art to the
present invention.
3. SUMMARY OF THE INVENTION
[0009] An ultrasonic or acoustic viscosity sensor 1 for measuring
fluid viscosity or fluid flow of a fluid in real time is provided.
In one embodiment, the viscosity sensor comprises:
[0010] an ultrasonic or acoustic actuator 10;
[0011] at least one microprobe 20 (also referred to herein as a
"probe") for sensing ultrasound induced motion; and
[0012] means for generating motion via coupling between the
actuator and the at least one microprobe.
[0013] In one embodiment, the ultrasonic or acoustic viscosity
sensor comprises a substrate.
[0014] In another embodiment, the ultrasonic actuator couples
motion to the at least one microprobe through the substrate.
[0015] In another embodiment, the actuator is driven at at least
one of its mechanical resonance frequencies.
[0016] In another embodiment, the at least one microprobe is driven
at at least one of its mechanical resonance frequencies by the
actuator.
[0017] In another embodiment, the actuator is a silicon ultrasonic
longitudinal mode actuator.
[0018] In another embodiment, the actuator is actuated by at least
one piezoelectric element.
[0019] In another embodiment, the piezoelectric element is
adhesively attached.
[0020] In another embodiment, the piezoelectric element is a thin
piezoelectric film.
[0021] In another embodiment, the silicon ultrasonic longitudinal
mode actuator has varying cross-section for amplifying motion
imparted to the at least one microprobe.
[0022] In another embodiment, the ultrasonic or acoustic viscosity
sensor operates in acoustic frequencies between 10 Hz to 20
kHz.
[0023] In another embodiment, the ultrasonic or acoustic viscosity
sensor operates in mid ultrasonic frequencies between 20 kHz to 500
kHz.
[0024] In another embodiment, the ultrasonic or acoustic viscosity
sensor operates in upper ultrasonic frequencies between 500 kHz to
10 MHz.
[0025] In another embodiment, the ultrasonic or acoustic viscosity
sensor generates high amplitude (e.g., 1-10 .mu.m) motion in the
fluid of interest.
[0026] In another embodiment, the ultrasonic or acoustic viscosity
sensor comprises means for measuring damping of the mechanical
vibrations of the at least one microprobe in the fluid of
interest.
[0027] In another embodiment, the means for measuring damping of
the mechanical vibrations comprises at least two strain gauges
(e.g., polysilicon strain gauges) or piezoresistors located at the
junction of the actuator 10 and the microprobes 20, wherein at
least one of the strain gauges or piezoresistors measures the
amplitude of motion of the at least one microprobe and wherein at
least one of the strain gauges or piezoresistors measures the
motion of the actuator.
[0028] In another embodiment, at least one of the strain gauges or
piezoresistors is a polysilicon resistor
[0029] In another embodiment, at least one of the strain gauges or
piezoresistors is integrated (or positioned) at the junction of the
actuator 10 (e.g., silicon horn actuator) and the microprobe.
[0030] In another embodiment, at least one of the strain gauges or
piezoresistors has a nominal resistance of about 10 k.OMEGA..
[0031] In another embodiment, the ultrasonic or acoustic viscosity
sensor comprises means 230 for measuring immersion depth of the
microprobe in the fluid.
[0032] In another embodiment, the means for measuring immersion
depth is based on measurement of capacitance between two
electrodes.
[0033] In another embodiment, the ultrasonic or acoustic viscosity
sensor comprises a capacitance-based immersion depth sensor.
[0034] In another embodiment, the capacitance-based immersion depth
sensor enables self-calibration of the viscosity sensor.
[0035] In another embodiment, the capacitance-based immersion depth
sensor comprises means for distance-coding capacitance.
[0036] In another embodiment, the means for distance-coding
capacitance is formed by at least two metal traces on the
capacitance-based immersion depth sensor.
[0037] In another embodiment, the capacitance is measured as a
function of frequency to extract the dielectric constant of the
fluid as a function of frequency.
[0038] In another embodiment, the ultrasonic or acoustic viscosity
sensor comprises computational circuitry functionally connected to
the microprobe.
[0039] In another embodiment, the ultrasonic or acoustic viscosity
sensor is integrated into a strip format.
[0040] In another embodiment, the strip format is a blood-glucose
measurement strip
[0041] In another embodiment, the ultrasonic or acoustic viscosity
sensor comprises:
[0042] a silicon horn 15;
[0043] a piezoelectric actuator element 16; and
[0044] a multi-sensor microprobe 20 for sensing ultrasound induced
motion,
wherein:
[0045] the multi-sensor multiprobe comprises: [0046] at least one
polysilicon strain gauge or piezoresistor 200 connected in a [0047]
Wheatstone bridge 250 configuration, and [0048] a capacitance-based
immersion depth sensor 230.
[0049] In a specific embodiment, the actuator comprises the silicon
horn 15 and the piezoelectric actuator element 16.
[0050] In one embodiment, the ultrasonic or acoustic viscosity
sensor is mid-frequency resonant, e.g., 20-500 kHz.
[0051] In another embodiment, the ultrasonic or acoustic viscosity
sensor is high-frequency resonant, e.g., 500 kHz-10 MHz.
[0052] In another embodiment, the ultrasonic or acoustic viscosity
sensor generates high amplitude motion. In a specific embodiment,
the amplitude of the high amplitude motion is 1 .mu.m or
greater.
[0053] In another embodiment, the ultrasonic or acoustic viscosity
sensor comprises at least one multi-sensor microprobe.
[0054] In another embodiment, the at least one multi-sensor
microprobe comprises a plurality of strain gauges or
piezoresistors.
[0055] In another embodiment, the microprobe comprises a plurality
of capacitance-based immersion depth sensors.
[0056] In another embodiment, the multi-sensor microprobe comprises
a plurality of capacitance-based immersion depth sensors.
[0057] A method for measuring fluid viscosity or fluid flow of a
fluid of interest is also provided. In one embodiment, the method
comprises the steps of:
[0058] providing the ultrasonic or acoustic viscosity sensor of
claim 1; and
[0059] measuring viscous damping of non-linear flexural vibrations
of the microprobe induced by the .lamda./2 longitudinal resonance
of the silicon horn. The fluid viscosity can then be calculated
using the viscous damping measurement(s) as described herein.
[0060] In one embodiment, the non-linear flexural vibrations of the
microprobe comprise longitudinal and flexural oscillating
motions.
[0061] In another embodiment, the step of measuring viscous damping
comprises measuring viscous damping at multiple frequencies.
[0062] In another embodiment, a first of the multiple frequencies
is the actuator longitudinal mode frequency and a second of the
multiple frequencies is the microprobe resonance frequency.
[0063] In another embodiment, the microprobe resonance is excited
parametrically at a frequency different then the actuation
frequency for simultaneous dual frequency actuation. shear
(non-linear vibration) damping and bulk (silicon horn's
longitudinal resonance) damping at two different frequencies,
thereby simultaneously measuring the static (bulk) viscosity and
dynamic (shear-rate dependent) viscosity of the fluid of
interest.
[0064] In another embodiment, the method comprises calculating the
frequency and magnitude of a signal corresponding to the
multi-frequency excitation from the Fast Fourier Transform (FFT) of
the amplified Wheatstone bridge voltage signal corresponding to the
motion of a cantilever of the at least one microprobe.
[0065] In another embodiment, the method comprises the step of
calibrating immersion depth. In a specific embodiment, the
calibrating step comprises measuring capacitance at discrete
steps.
[0066] In another embodiment, the method comprises the step of
determining the shear decay length (.delta.) at the actuator
resonance, wherein .delta. is given by
.delta. = ( 2 .eta. .omega..rho. ) 1 2 ##EQU00001##
[0067] where .eta. and .rho. are the fluid viscosity and density,
respectively, and .omega.=2.pi.f is the angular frequency of
flexural vibration of the at least one microprobe
[0068] In another embodiment, the method comprises the step of
measuring the viscosity of the fluid of interest at different shear
rates. In one embodiment, this step comprises varying piezoelectric
transducer (PZT) actuation voltage.
[0069] In another embodiment, the method comprises the step of
assessing coagulation time of the fluid of interest (e.g.,
blood).
[0070] In another embodiment, the method comprises the step of
measuring the coagulation of the fluid of interest (e.g.,
blood).
4. BRIEF DESCRIPTION OF THE DRAWINGS
[0071] The present invention is described herein with reference to
the accompanying drawings, in which similar reference characters
denote similar elements throughout the several views. It is to be
understood that in some instances, various aspects of the invention
may be shown exaggerated or enlarged to facilitate an understanding
of the invention.
[0072] FIG. 1. An embodiment of the viscosity sensor 1. Top left:
Photograph of the viscosity sensor 1. Piezoresistors 200 (in this
embodiment, .about.9 k.OMEGA.) are located at the junction of an
actuator 10 (in this embodiment, a catenoid silicon horn 10) and a
pair of microprobes 20 to sense ultrasound-induced strain
oscillations. PZT plates 40 (piezoresistive transducer plates). Top
right: metal traces 230 (in this embodiment, 10 .mu.m.times.0.2
.mu.m platinum traces) are located such that there is a capacitor
formed between the two wires. The two wires consist of discrete
steps along the microprobes, for provided stepped capacitance
change along the length of the probe. As the microprobe 20 is
inserted in a liquid, the different dielectric constant of the
liquid results in increase in capacitance as the probe immersion
depth is increased. As the liquid crosses each of the stepped
locations, the capacitance change is stepped as well. Since the
distance between the steps is known, the position and velocity of
insertion can be calibrated from the change in capacitance curves.
A second capacitance sensor can be incorporated opposite to the
sensor on the opposite side of the probes to provide a differential
capacitance measurement. Incorporating multiple electrodes on
multiple probes can lead to determination of the meniscus level at
each sensor, enabling the measurement of the tilt of the probes
with respect to the fluid meniscus. In this embodiment, two
polysilicon resistors are placed on the actuator side, and two
polysilicon resistors are placed on the two probes. The four
resistors 200 are connected in a Wheatstone bridge 250
configuration such that change in the resistance due to strain in
the probes is measured differentially compared to the strain on the
actuator side. Bottom: The microprobes can be used for
multi-sensing viscosity (bulk and shear) measurements in
microdroplet fluid samples at specific or desired immersion depths.
The diagram depicts simultaneous detection of shear-based viscosity
at f.sub.1, for which the shear viscous depth of a few microns is
much smaller than the droplet diameter and bulk viscosity at
f.sub.2 where the wavelength is typically much larger than the
droplet diameter. Custom PC-board 180.
[0073] FIGS. 2A-B. A. Frequency spectrum of Wheatstone bridge
voltage showing the longitudinal and flexural vibration. B.
Two-dimensional laser Doppler vibrometer measurement of
microprobe's flexural displacement and surface plot along the
length of the microprobes indicating the mode shape with three
nodes. The 2-D scan is of z-displacement velocity of microprobe
flexural vibration along the probe's length. Two piezo-electric
transducers (14.2.times.5 mm) are driven at a longitudinal
resonance of the silicon horn (f=109.8 KHz). The surface plot shows
the out-of-plane flexural vibration of the microprobes at 30.4
KHz.
[0074] FIG. 3. Capacitance-based calibration of microprobe
immersion depth in physiological saline (0.9% w/v NaCl). Platinum
traces (10 .mu.m wide and 0.2 .mu.m thick) have discrete steps of
0.75 mm and capacitance is measured at every step.
[0075] FIG. 4. Amplitude of microprobe flexural vibration in
de-ionized (DI) water when compared to air, for increasing PZT
actuation voltages at varying immersion depths of microprobe (1-3
mm).
[0076] FIG. 5. Damping of microprobe vibration in ethylene glycol
solutions (2 mm immersed) of varying viscosities at 8V.sub.pp PZT
actuation. The viscous damping of the microprobe is modeled
assuming the immersed part to be an oscillating sphere.
[0077] FIG. 6. Frequency variation of microprobe flexural vibration
in whole rat blood (.about.1 mm immersed) at 10V.sub.pp PZT
actuation. Phase I: The red blood cells (RBCs) coat the microprobe
within the shear thickness (.about.4 .mu.m) and load the vibration
frequency reducing it. Phase II: The blood starts to coagulate with
the RBCs clotting with clot sizes bigger than the shear thickness,
thus lessening the loading on the vibration (increasing the
frequency). Phase III: End of blood coagulation with the formation
of fibrous structures and large clots--the residual damping is due
to few RBCs, the plasma and the fibrous network of clots.
[0078] FIG. 7. Another embodiment of the viscosity sensor 1. In
this embodiment, the viscosity sensor 1 comprising the microprobes
20 is integrated into a typical strip test, much like those used
for blood-glucose measurement. Commercially available glucose
strips have a small wicking channel formed by two plastic or paper
laminates such the blood is wicked into the strips. Electrical
interconnects 150 are provided at the back end of the viscosity
sensor 1 for electrochemical measurements. The electrical
interconnects 150 can plug into a reader for performing the
measurements. The blood enters the chamber with the microprobes 20
through the wick 160. Once the blood is collected, simultaneous
measurements of glucose levels and viscosity, and thus PT and PTT
coagulation times can be performed. As the blood constituents
change as blood plasma evaporates, the levels of the liquid on the
probes can be measured continuously to maintain accuracy of
viscosity measurement.
[0079] FIG. 8. Blood coagulation tests. Diagram of two
hemorheological tests that are blood coagulation tests and that are
currently used for diagnosis and monitoring of diseases.
Prothrombin Time (PT) test. Partial Thromboplastic Therapy (PTT)
test.
[0080] FIGS. 9A-B. A. One embodiment of the invention showing
measurement of fluid flow inside a channel or tunnel 300 by
piezoresistive probes 20. B. .about.2 mm of the probe tips is
inside the tunnel/channel 300. Tubule 400. Syringe driver 600. See
Section 6.2 for details.
[0081] FIG. 10. A voltage measurement taken when a channel is
filling with water during the experiment described in Section 6.2.
System settles in 3.5 sec.
[0082] FIG. 11. A voltage measurement taken when a channel is
filling with water during the experiment described in Section 6.2.
Measurement started after system settlement.
[0083] FIG. 12. A voltage measurement taken when the water depth in
the channel is constant during the experiment described in Section
6.2.
[0084] FIGS. 13A-B. A. Another embodiment of the viscosity sensor
in which the resonant probe is directly used to puncture skin,
extract blood and measure the blood level on the probes, while
measuring the viscosity of the blood. B. The capacitors formed by
electrodes on the probes can measure the level of the fluid on the
probes. Two different probes can measure the meniscus at two
different points enabling the measurement of angle of the meniscus.
This will enable a way to measure the tilt of the probes with
respect to the meniscus for further improving the accuracy of the
viscosity measurement.
5. DETAILED DESCRIPTION OF THE INVENTION
[0085] For clarity of disclosure, and not by way of limitation, the
detailed description of the invention is divided into the
subsections set forth below.
[0086] 5.1. Ultrasonic Viscosity Sensor
[0087] An ultrasonic or acoustic viscosity sensor (also referred to
herein as a viscometer) is provided that can be used to accurately
measure viscosity for fluid samples of less than 1 .mu.l in volume.
Methods for measuring viscosity for fluid samples of less than 1
.mu.l in volume are also provided. The viscosity sensor and methods
based thereon enable simultaneous measurement of bulk and dynamic
(shear-rate dependent) viscosity of a non-Newtonian fluid. The
viscosity of any fluid known in the art can be measured. For
example, the viscosity of any known bodily fluid, such as blood,
pus, mucous, semen, vaginal secretion, saliva or tears, can be
indicative of undesired biochemical reactions owing to the presence
of disease or foreign agents. In a specific embodiment, the fluid
is blood.
[0088] Bulk and dynamic viscosity of the non-Newtonian fluid can be
measured simultaneously without separating constituents of the
fluid, and thus distinguishing the effect of constituents on the
viscosity. Dynamic viscosity of the non-Newtonian fluid can be
estimated at varying shear rates, to study the deformability of the
constituents of the fluid as a function of shear rate. The
viscosity sensor can also be tuned to measure viscosity of a
non-Newtonian fluid based on the particle size of the fluid's
constituents.
[0089] In one embodiment, the ultrasonic or acoustic viscosity
sensor measures fluid viscosity or fluid flow of a fluid in real
time. The viscosity sensor can comprise:
[0090] an ultrasonic or acoustic actuator;
[0091] at least one microprobe for sensing ultrasound induced
motion of the at least one microprobe; and
[0092] means for generating motion via coupling between the
actuator and the at least one microprobe.
[0093] Actuator
[0094] The viscosity sensor can comprise an ultrasonic or acoustic
actuator. In one embodiment, the actuator is driven at its
mechanical resonance frequencies.
[0095] In another embodiment, the actuator drives at least one
microprobe at its mechanical resonance frequency.
[0096] In another embodiment, the actuator is a silicon ultrasonic
longitudinal mode actuator (FIG. 1) (also referred to herein as a
"silicon horn" actuator). Silicon horns and silicon horn actuators
are well known in the art, see, e.g., A. Ramkumar, X. Chen and A.
Lal, Silicon Ultrasonic Horn Driven Microprobe Viscometer,
Proceedings of the International IEEE Ultrasonics Symposium (UFFC
2006), Vancouver, CANADA, October 2006, pp. 1449-1452; Chen, X.,
Lal, A., Riccio, M., and Gilmour Jr., R., Ultrasonically Actuated
Silicon Microprobes for Cardiac Signal Recording, IEEE Transactions
on Biomedical Engineering, 53 (8) 2006, p. 1665; Chen, X., Lal, A.,
Integrated Pressure and Flow Sensor in Silicon-based Ultrasonic
Surgical Actuator, IEEE Ultrasonics, Ferroelectrics, and Frequency
Control Society Symposium, 2001, Atlanta; Chen, X., Lal, A.,
Micromachined Ultrasonic Ophthalmic Microsurgical tool with
integrated Pressure sensor, Digest of Technical publications,
International Conference on Solid State Sensors and Actuators,
2001, Munich, pp. 424-427; Son, I., Lal, A., Hubbard, B., Olsen,
T., A multifunctional silicon-based microscale surgical system,
Sensors and Actuators A, Volume 91, Issue 3, pp. 351-356, 2000;
Lal, A., White, R. M., Silicon micro-fabricated horns for power
ultrasonics, Sensors and Actuators, vol. A54, pp. 542-546, 1996;
Amit Lal, Micromachined Silicon Ultrasonic Longitudinal Mode
Actuators: Theory and Applications to Surgery, Pumping, and
Atomization, Ph.D. Dissertation, University of California,
Berkeley, 1996, pp. 1-175; U.S. Pat. No. 5,728,089 to Lal et al.,
Mar. 17, 1998. Such an actuator can be micromachined using methods
known in the art.
[0097] The ultrasonic actuator can couple motion to the at least
one microprobe through a substrate. Suitable substrates for
coupling such motion are known in the art (see, e.g., WO2008/151328
(PCT/US2008/066375; U.S. Pat. No. 5,728,089 to Lal et al., Mar. 17,
1998). The substrate can be made from a semiconducting material.
The substrate can comprise both a body portion and the horn portion
(also referred to herein as a "catenoid" or catenoid-shaped horn)
projecting forward from the body portion. The horn portion can have
a blade portion that is free to vibrate.
[0098] In another embodiment, silicon ultrasonic longitudinal mode
actuator has a varying cross-section for amplifying motion imparted
to the at least one microprobe.
[0099] In another embodiment, the blade portion can have a forward
edge that is significantly thinner than the body portion.
[0100] In another embodiment, the actuator imparts mechanical
energy to blade portion of the horn. The actuator can be
mechanically coupled to the surface (e.g., the top surface) of the
body portion.
[0101] The viscosity sensor can also comprise piezoelectric
elements art (see, e.g., WO2008/151328 (PCT/US2008/066375; U.S.
Pat. No. 5,728,089 to Lal et al., Mar. 17, 1998). The ultrasonic
actuator can be actuated by piezoelectric elements. In one
embodiment, the piezoelectric elements are adhesively attached to
the actuator. In another embodiment, the piezoelectric elements are
thin piezoelectric films.
[0102] Microprobes
[0103] The ultrasonic viscosity sensor comprises one or more
microprobes. Suitable microprobes for use in the viscosity sensor
and the methods disclosed herein are described in WO2008/151328
(PCT/US2008/066375). The design and fabrication of the microprobe
can be accomplished using methods of design and fabrication known
in the art for other ultrasonic microprobes (see, e.g.,
WO2008/151328; A. Ramkumar, X. Chen and A. Lal, IEEE Ultrasonics
Symposium 2006, pp. 1449-1452, 2006).
[0104] In one embodiment, the microprobe is formed from silicon.
Other less expensive, mass-produced materials such as stainless
steel, nickel, and others known in the art can be used to reduce
the cost of making the microprobe. Suitable dimensions for the
microprobe can be easily determined by the skilled artisan. In the
embodiment shown in FIG. 1, the silicon microprobe 20 is 5 mm long,
100 .mu.m long wide and 140 .mu.m thick projecting outwards at the
tip of a silicon horn 10.
[0105] The microprobe 20 can comprise a probe body that is formed
from silicon. Designs for suitable microprobes for use in the
invention are known in the art (see, e.g., WO2008/151328
(PCT/US2008/066375). In one embodiment, the probe body can comprise
a core; a lower blade that is integrally formed with the core and
that terminates in a lower blade pointed tip extending beyond a
distal end of the core; and an upper blade that is formed from
silicon nitride deposited on an upper surface of the core opposite
the lower blade; terminates in an upper blade pointed tip extending
beyond the distal end of the core; and comprises a strain gauge or
polysilicon resistor that outputs a signal indicative of strain in
the upper blade. The strain gauge can comprise elements etched on
the upper blade. The strain gauge can comprise a Wheatstone bridge
250 (for details, see WO2008/151328 (PCT/US2008/066375).
[0106] The microprobe can further comprise a base formed from
silicon and having a proximal end and a distal end, the probe body
extending from the distal end. The base can taper from the proximal
end to the distal end. In a preferred embodiment, the base tapers
with a catenoid shape. The microprobe can further comprise a second
probe body extending from the base (for details, see WO2008/151328
(PCT/US2008/066375).
[0107] In one embodiment, the strain gauge or piezoresistor 200 can
comprise two sensing resistors 220 so positioned on the probe body
as to vary in resistance with strain in the upper blade, and two
constant resistors 210 so positioned on the body as not to vary in
resistance with strain in the upper blade.
[0108] In a specific embodiment, the microprobe defines a channel
or tunnel 300 dimensioned for fluid flow (for details, see
WO2008/151328 (PCT/US2008/066375).
[0109] In one embodiment, the viscosity sensor comprises a single
multi-sensor microprobe integrated on the silicon ultrasonic
longitudinal mode actuator (silicon horn). In another embodiment,
the viscosity sensor can comprise a plurality (e.g., dual) of
multi-sensor microprobes integrated on the silicon ultrasonic
longitudinal mode actuator (silicon horn). Such an arrangement can
increase the sampling of viscosity measurement.
[0110] In another embodiment, the multi-sensor microprobe comprises
a probe body. In another embodiment, the multi-sensor microprobe
comprises two probe bodies extending from the body. In one
embodiment, the probe bodies are separated from one another by a
distance of preferably no more than 1 mm.
[0111] In another embodiment, the multi-sensor microprobe comprises
two or more sensors, each of which is coupled to a respective probe
body to sense a property of that probe body. Such designs for
multi-sensor microprobes are known in the art (for details, see
WO2008/151328 (PCT/US2008/066375)). In a specific embodiment, the
sensors are a plurality of polysilicon strain gauges or
piezoresistors and/or capacitance-based immersion depth
sensors.
[0112] Single-sensor and multi-sensor microprobes can also
comprise, or be functionally connected to, computational circuitry
that is responsive to the sensed properties and programmed with
instructions to recognize a contact event sensed by one of the
probe bodies; to recognize common mode noise sensed by the other of
the probe bodies; and output an output signal indicative of the
contact event minus the common mode noise (see, e.g., WO2008/151328
(PCT/US2008/066375)). In a specific embodiment, the probe bodies
are separated from one another by a distance of no more than 500
.mu.m. The design of such computational circuitry is known in the
art
[0113] In one embodiment, the fluid sample can be ultrasonically
driven in the ultrasonic viscosity sensor, e.g., by capillary
ultrasonic drive and bubble ultrasonic drive of fluid samples.
[0114] The microprobes can be driven by different modalities, e.g.,
piezoelectrically driven microprobes, ultrasonically driven
microprobes (e.g., glass capillaries), or acoustically driven
microprobes (e.g., driven by acoustically created bubbles).
[0115] Piezoelectrically driven probes can be configured using
methods known in the art. Various configurations to couple
ultrasonic energy into microprobes, whether directly by integrated
piezoelectric or by remote ultrasonic actuators can be used.
Piezoresistive readouts can be integrated into the viscosity
sensor.
[0116] In another embodiment, the viscosity sensor can comprise a
piezoelectrically driven glass capillary.
[0117] In another embodiment, the viscosity sensor can comprise an
optical imager.
[0118] In another embodiment, fluid can be bubble-driven in the
viscosity sensor. Electrolytically created, but acoustically driven
microbubbles in samples can be used to measure viscosity.
[0119] In another embodiment, the viscosity sensor can comprise a
multi-sensor microprobe. The microprobe can comprise, for example,
one or more strain gauges and/or immersion depth capacitance
sensors. In other embodiments, the viscosity sensor comprises a
single sensor microprobe.
[0120] 5.2. Operation of the Viscosity Sensor
[0121] In another embodiment, the viscosity sensor operates in
acoustic frequencies between 10 Hz to 20 kHz.
[0122] In another embodiment, the viscosity sensor operates in mid
ultrasonic frequencies between 20 kHz to 500 kHz.
[0123] In another embodiment, the viscosity sensor operates in in
upper ultrasonic frequencies between 500 kHz to 10 MHz.
[0124] In another embodiment, the viscosity sensor generates high
amplitude (e.g., 1-10 .mu.m motion in the fluid of interest.
[0125] In certain embodiments, the ultrasonic or acoustic viscosity
sensor can be used to measure viscosity in small volumes of a fluid
of interest, e.g., less than 1 .mu.l. In other embodiments, the
sensor can measure viscosity in sample volumes that are 1-10 .mu.l,
10-20 .mu.l, 20-30 .mu.l, 30-40 .mu.l, 40-50 .mu.l, 50-60 .mu.l,
60-70 .mu.l, 70-80 .mu.l, 80-90 .mu.l or 90-100 .mu.l or greater
than 100 .mu.l.
[0126] In a specific embodiment, the viscosity sensor uses a low
sample volume. In one embodiment, the dimensions of the microprobe
can accommodate sample volumes 1 .mu.l or lower. This sample size
corresponds, for example, to a blood sample associated with
diabetes in-home sugar testing equipment.
[0127] In one embodiment, the ultrasonic or acoustic viscosity
sensor is a mid-frequency (30-50 kHz) resonant viscosity
sensor.
[0128] In another embodiment, the sensor generates high amplitude
motion that is 1 .mu.m or greater.
[0129] The sensor generates the high amplitude (>1 .mu.m) motion
via nonlinear coupling between longitudinal to transverse motion to
measure fluid viscosity.
[0130] Vibrations are induced by the .lamda./2 longitudinal
resonance of the ultrasonic longitudinal mode actuator (silicon
horn). Since shear (non-linear vibration) and bulk (silicon horn's
.lamda./2 longitudinal resonance) damping can be observed at two
different frequencies, the viscosity sensor can simultaneously
measure the static (bulk) and dynamic (shear-rate dependent)
viscosity of a fluid (FIG. 1, bottom diagram).
[0131] In another embodiment, the ultrasonic actuator can be a
remote ultrasonic actuator.
[0132] FIG. 1 shows a specific embodiment of the viscosity sensor.
FIG. 1, top left is a photograph of the viscosity sensor.
Piezoresistors 200 (in this embodiment, .about.9 k.OMEGA.) are
located at the junction of an actuator (in this embodiment, a
catenoid silicon horn) and a pair of microprobes to sense
ultrasound-induced strain oscillations. PZT plates (piezoresistive
transducer plates). In the embodiment shown in FIG. 1, four
polyresistors 200, consisting of two constant polyresistors 210 and
two sense polyresistors 220 are connected in a Wheatstone bridge
250 configuration.
[0133] Furthermore, since fluid viscosity is a strong function of
probe insertion depth, the viscosity sensor can comprise means for
measuring and/or calibrating immersion depth. Such means can be
used for self-calibration of the viscosity sensor. In one
embodiment, the means for measuring immersion depth is based on
measurement of capacitance between two electrodes.
[0134] In a specific embodiment, the means for measuring immersion
depth are capacitance-based immersion depth sensors that can be
used for self-calibration of the viscosity sensor, i.e., for
calibrating a viscosity measurement. The viscosity sensor thus can
measure the vibration damping at varying microprobe immersion
depths.
[0135] The capacitance-based immersion depth sensor can comprise
distance-coding capacitance. In another embodiment, the
distance-coding capacitance is formed by at least two metal traces
on the capacitance-based immersion depth sensor. In another
embodiment, the capacitance is measured as a function of frequency
to extract the dielectric constant of the fluid as a function of
frequency.
[0136] In one embodiment, immersion depth calibration is
implemented by measuring the capacitance at discrete steps (FIG.
1), which provide distinct motion artifacts from velocity and depth
measurement. FIG. 1, bottom, is a diagram showing the microprobes
being used for multi-sensing viscosity (bulk and shear)
measurements in microdroplet fluid sample. The multi-sensing
measurements can be conducted at specific or desired immersion
depths. The diagram depicts simultaneous detection of shear-based
viscosity f.sub.1 and bulk viscosity at f.sub.2 in low sample
volumes because of vibration wavelengths being much smaller than
droplet diameter.
[0137] For example, the device can comprise means for
distance-coding of a capacitance measurement. In one embodiment,
means for distance-coding of capacitance can comprise a plurality
of platinum traces (e.g., 10 .mu.m wide and 0.2 .mu.m thick)
positioned on the device. Such means can be, in certain
embodiments, co-fabricated on the microprobes.
[0138] FIG. 3 shows the linear dependence of the capacitance signal
with increasing depth of immersion of the microprobes in the
liquid
[0139] FIG. 1, top right, shows metal traces (in this embodiment,
10 .mu.m.times.0.2 .mu.m platinum traces) that are located at
discrete steps along the microprobes that serve as
capacitance-based probe immersion depth sensors. The spacing of the
sensors can be readily determined by the skilled artisan.
[0140] The ultrasonic or acoustic viscosity sensor can comprise
means for measuring damping of the mechanical vibrations of the at
least one microprobe in the fluid of interest.
[0141] In one embodiment, the means for measuring damping of the
mechanical vibrations comprises at least two piezoresistors or
strain gauges located at the junction of the silicon horn and the
microprobes, wherein at least one of the piezoresistors measures
the amplitude of motion of the at least one microprobe and wherein
at least one of the piezoresistors measures the motion of the
actuator.
[0142] The resistance of the piezoresistor or strain gauge changes
with the strain experienced. In one embodiment, the piezoresistor
or strain gauge has a nominal resistance of about 10 k.OMEGA.. In
other embodiments, the piezoresistor or strain gauge has a nominal
resistance of 1-5, 5-10, 10-15 or 15-20 or 20-50 k.OMEGA.. Other
suitable resistances can be easily determined by the skilled
artisan.
[0143] The piezoresistor or strain gauge can be integrated at the
junction of the actuator and the microprobe to measure longitudinal
and flexural oscillating motion (see FIGS. 1 and 2). In a specific
embodiment, the strain gauge is a polysilicon strain gauge or
polysilicon resistor (polyresistor) (FIG. 1).
[0144] In another embodiment, the resistors or strain gauges are
connected in a Wheatstone bridge 250 configuration (FIG. 1, see,
e.g., A. Ramkumar, X. Chen and A. Lal, IEEE Ultrasonics Symposium
2006, pp. 1449-1452, 2006).
[0145] The silicon horn actuator can be actuated at its .lamda./2
longitudinal resonance, which exerts longitudinal strain at the tip
of the horn and induces a parametrically excited flexural vibration
mode in the microprobe. The magnitude of signal corresponding to
both the oscillation frequencies can be measured from the
Fast-Fourier Transform (FFT) of the amplified Wheatstone bridge
voltage.
[0146] FIGS. 2A-B show a two-dimensional (2D) laser Doppler
vibrometer scan of silicon horn actuator-based microprobes that
indicates a mode with three displacement nodes.
[0147] In specific embodiments, mid- or high-frequency resonant
ultrasonic viscosity sensors for measuring fluid viscosity or flow
are provided (FIG. 1). The mid- or high-frequency resonant
ultrasonic viscosity sensor can comprise:
[0148] a silicon horn 10; and
[0149] a multi-sensor microprobe 20 for sensing ultrasound induced
strain oscillations, wherein:
[0150] the multi-sensor multiprobe comprises: [0151] at least one
strain gauge (e.g. polysilicon strain gauge) or piezoresistor 200
connected in a Wheatstone bridge 250 configuration, and [0152] a
capacitance-based immersion depth sensor 230, and
[0153] the sensor generates high amplitude motion via nonlinear
coupling between longitudinal to transverse motion.
[0154] In one embodiment, the viscosity sensor is mid-frequency
resonant and the mid-frequency is 20-500 kHz.
[0155] In another embodiment, the viscosity sensor is
high-frequency resonant and wherein the high-frequency is 500
kHz-10 MHz.
[0156] 5.3. Methods for Measuring Fluid Viscosity Using the
Ultrasonic Viscosity Sensor
[0157] Methods for measuring fluid viscosity are also provided. In
one embodiment, the method can comprise the steps of providing an
ultrasonic viscosity sensor and monitoring viscous damping of
non-linear flexural vibrations of microprobes on the viscosity
sensor.
[0158] In one embodiment, the method for measuring fluid viscosity
or fluid flow of a fluid of interest comprising the steps of:
[0159] providing an ultrasonic viscosity sensor as described
herein; and
[0160] measuring viscous damping of non-linear flexural vibrations
of the microprobe induced by the .lamda./2 longitudinal resonance
of the silicon horn.
[0161] In one embodiment, non-linear flexural vibrations of the
microprobe can comprise longitudinal and flexural oscillating
motions.
[0162] In another embodiment, the step of measuring viscous damping
comprises measuring viscous damping at multiple frequencies. In
another embodiment, a first of the multiple frequencies is the
actuator longitudinal mode frequency and a second of the multiple
frequencies is the microprobe resonance frequency.
[0163] In another embodiment, the step of measuring viscous damping
can comprise measuring shear (non-linear vibration) damping and
bulk (silicon horn's .lamda./2 longitudinal resonance) damping at
two different frequencies, thereby simultaneously measuring the
static (bulk) viscosity and dynamic (shear-rate dependent)
viscosity of the fluid of interest.
[0164] In another embodiment, the method can comprise the step of
calculating the frequency and magnitude of a signal corresponding
to the multi-frequency excitation from the Fast Fourier Transform
(FFT) of the amplified Wheatstone bridge voltage signal
corresponding to the motion of a cantilever of the microprobe.
[0165] In another embodiment, the method can comprise the step of
calibrating immersion depth. In one embodiment, the calibrating
step can comprise measuring capacitance at discrete steps.
[0166] As mentioned above, the method comprises measuring viscous
damping of non-linear flexural vibrations of the microprobe induced
by the .lamda./2 longitudinal resonance of the silicon horn.
Damping of the microprobe occurs because the liquid on the surface
of the microprobe is viscously entrained with the shear acoustic
wave penetrating only a small distance into the fluid sample. This
decay length (.delta.) is given by
.delta. = ( 2 .eta. .omega..rho. ) 1 2 ##EQU00002##
(I. Etchart, H. Chen, P. Dryden, J. Jundt, C. Harrison, K. Hsu, F.
Marty, B. Mercier, "MEMS sensor for density-viscosity sensing in a
low-flow microfluidic environment", Sensors and Actuators A., vol.
141, pp. 266-275, 2008), where .eta. and .rho. are the fluid
viscosity and density, respectively, and .omega.=2.pi.f is the
angular frequency of flexural vibration of the microprobe(s).
[0167] The viscosity of the fluid of interest can be estimated at
different shear rates by varying the piezoelectric transducer (PZT)
actuation voltage. In one embodiment, the viscosity of the fluid of
interest is used to assess coagulation time. In another embodiment,
the coagulation progression of the fluid of interest is measured.
In a specific embodiment, the coagulation progression of blood is
measured.
[0168] For example, to perform such a calculation for blood
viscosity, based on known normal blood plasma, .rho.=1060
Kg/m.sup.3 and .eta.=0.001 Poise, the maximum decay length will be
.delta..sub.0.about.2.1 .mu.m at the silicon horn's .lamda./2
longitudinal resonance (109.8 KHz). This length is much smaller
than the dimensions of the red blood cells (RBCs), which are on the
order of 8 .mu.m in length. Unless the cells are very close the
surface of the microprobes, the mechanical movement that entrains
the fluid will not perturb the RBCs, thus, the microprobe only
"sees" the plasma surrounding the cells and measures plasma
properties (K. K. Kanazawa, J. G. Gordon, "Frequency of a quartz
microbalance in contact with liquid", Anal. Chem., vol. 57, no. 8,
pp. 1770-1771, 1985). The advantage of this approach is that it
allows the extraction of information on plasma viscosity without
having to separate the plasma from the whole blood.
[0169] The frequency of lower parametric flexural vibrations
manifested in the microprobes can be adjusted so that the decay
length is greater than 8 .mu.m, thus allowing for whole blood
viscosity measurement with the effect of RBCs. By measuring both
whole blood and plasma viscosity simultaneously, the effect of RBCs
can be precisely determined. Also, in embodiments in which the
viscosity sensor is calibrated for different immersion depths, the
viscosity of a fluid such as blood can be estimated at different
shear rates by varying the piezoelectric transducer (PZT) actuation
voltage. It will be apparent to the skilled artisan that the
above-described approach can be used to extract information about
the viscosity of other fluids of interest.
[0170] 5.4. Uses for the Viscosity Sensor
[0171] The viscosity of any fluid of interest can be measured using
the viscosity sensor and methods based thereon.
[0172] In certain embodiments, the sensor can be used in clinical
and home-use devices that utilize viscosity data about fluids,
e.g., body fluids.
[0173] In one embodiment, the viscosity sensor is re-usable. For
example, the device can be cleaned in bleach after every
measurement and re-used in the clinical diagnostic setting for
maximal use of each sensor.
[0174] In another embodiment, several assays on different samples
can be performed simultaneously. For example, a glass capillary
could also be used to set up groups of data performing different
assays simultaneously.
[0175] In another embodiment, the viscosity sensor can be
integrated as a component into a hand-held reader, e.g., a reader
for blood-glucose measurements.
[0176] In another embodiment, the viscosity sensor can comprise a
plastic packaged capillary wick into which the fluid sample is
introduced or driven. The fluid sample can then be pumped into a
MEMS sensor of any suitable type known in the art.
[0177] In certain embodiments, the viscosity sensor is used in
clinical settings as a probe can be easily inserted in microliters
of sample collected from patients, such as a drop of blood. In one
embodiment, the viscosity sensor is integrated into a strip format.
In a specific embodiment, the strip format is a blood-glucose
measurement strip.
[0178] The device can be scaled to very small scales so that it can
be incorporated within existing blood-sensing strips, for easy
dissemination to millions of patients who already use blood-sugar
strips.
[0179] In a specific embodiment, the viscosity sensor is used to
perform a Prothrombin Time (PT) test. The PT test is an important
index for the activity of coagulation factors of the extrinsic
pathway; it is the coagulation time when tissue thromboplastin (a
tissue factor) and calcium ion are added into the plasma specimen
to induce coagulation formation (H. L. Bandey, R. W. Cemosek, W. E.
Lee III, L. E. Ondrovic, "Blood rheological characterization using
the thickness-shear mode resonator", Biosensors and Bioelectronics,
vol. 19, pp. 1657-1665, 2004).
[0180] In another embodiment, the viscosity sensor is used to
perform a Partial Thromboplastin Time (PTT) test. The PTT test is
an indicator of coagulation factors of the intrinsic pathway,
measuring the time whole blood takes to coagulate. PTT is often
used as a starting place when investigating the cause of a bleeding
or thrombotic episode. The PTT test is also used to monitor these
therapies. It does not directly measure the anticoagulants used but
measures their effect on blood clotting.
[0181] Since the viscosity sensor is capable of measuring plasma
viscosity and blood viscosity simultaneously, the viscosity sensor
can perform the PT and PTT coagulation tests simultaneously, thus
allowing for standardized results for both tests under the same
conditions, giving the doctor entire information about the blood
coagulation cascade.
[0182] In another embodiment of the viscosity sensor, shown in
FIGS. 13A-B, the microprobe is used directly to puncture skin,
extract blood and measure the blood level on the probes, while
measuring the viscosity of the blood (FIG. 13A). The capacitors
formed by electrodes on the probes can measure the level of the
fluid on the probes (FIG. 13B). Two different probes can measure
the meniscus at two different points enabling the measurement of
angle of the meniscus. This enables a way to measure the tilt of
the probes with respect to the meniscus for further improving the
accuracy of the viscosity measurement.
[0183] The following examples are offered by way of illustration
and not by way of limitation.
6. EXAMPLES
Example 1
Viscosity Measurement of Blood Samples
[0184] Introduction
[0185] This example demonstrates the development of a viscosity
sensor that can be used to accurately measure viscosity for less
than 1 .mu.l of blood sample, so that clinical and home-use devices
that utilize viscosity data can be realized. The device can be
integrated into a typical strip test, much like those used for
blood-glucose measurement (FIG. 7).
[0186] Medical diagnosis of disease often requires the viscosity of
body fluids, most importantly that of blood (R. S. Rosenson, A.
McCormick, E. F. Uretz, "Distribution of blood viscosity values and
biochemical correlates in healthy adults," Clin. Chem., vol. 42,
no. 8, pp. 1189-1195, 1996). Blood rheology in vivo can be used for
prediction of blood flow through capillaries of varying sizes,
while its in vitro rheology can be used for diagnosis and
monitoring of diseases. Also, changes in blood rheology have been
seen to contribute to or aggravate cardiovascular disorders such as
myocardial infarction and hypertension (S. Chien, "Blood rheology
in myocardial infarction and hypertension," Biorheology, vol. 23,
no. 6, pp. 633-653. 1986; T. Somer, H. J. Meiselman, "Disorders of
Blood Viscosity," Annals of Medicine, vol. 25, pp. 31-39, 1993).
Clinical hemorheological tests measure the red blood cell (RBC or
erythrocyte) aggregation by the erythrocyte sedimentation rate
(ESR) and the plasma viscosity. Physiological conditions leading to
alternation of plasma proteins (such as fibrinogen) results in a
change in the ESR and plasma viscosity. Correlation to blood
rheology can be used as a good indicator for prognosis, diagnosis
and/or monitoring disease processes such as cancer, cardiovascular
disease, etc. For example, inflammatory biomarkers (fibrinogen) and
plasma viscosity have been observed to be significantly associated
with death due to myocardial infarction or coronary heart disease
(S. G. Wanamethee, P. H. Whincup, A. G. Shaper, A. Rumley, L.
Lennon, G. D. O. Lowe," Circulating inflammatory and hemostatic
biomarkers are associated with risk of myocardial infarction and
coronary death, but not angina pectoris, in older men," J. Thromb.
Haemost, vol. 7, pp. 1605-1611, 2009). Incidence of type 2-Diabetes
has been shown to be strongly related to high whole blood viscosity
(WBV) (L. J. Tamariz, J. H. Young, J. S. Pankow, H.-C. Yeh, M. I.
Schmidt, B. Astor, F. L. Brancati, "Blood Viscosity and Hematocrit
as Risk Factors for Type 2 Diabetes Mellitus The Atherosclerosis
Risk in Communities (ARIC) Study," Am J. Epidemiol., vol. 168, no.
10, pp. 1153-1160, 2008), thus allowing a measure of blood
viscosity to be a prognosis tool. Also, monitoring plasma viscosity
can help in the prediction and diagnosis of plasma hyperviscosity
syndrome and sickle cell disease (T. Somer, H. J. Meiselman,
"Disorders of Blood Viscosity," Annals of Medicine, vol. 25, pp.
31-39, 1993). Additionally, a measure of blood viscosity can help
in monitoring interventions to alter blood viscosity such as diet,
alcohol, pharmaceuticals etc.
[0187] FIG. 8 shows the currently used hemorheological tests for
diagnosis and monitoring of diseases are blood coagulation tests
namely, Prothrombin Time (PT) and Partial Thromboplastic Therapy
(PTT). The Prothrombin Time (PT) test is an important index for the
activity of coagulation factors of the extrinsic pathway--it is the
coagulation time when tissue thromboplastin (a tissue factor) and
calcium ion are added into the plasma specimen to induce
coagulation formation. Warfarin and Coumarin are prescribed to slow
down the extrinsic pathway and the effectiveness is measured by the
PT test.
[0188] The Partial Thromboplastin Time (PTT) test is an indicator
of coagulation factors of the intrinsic pathway, measuring the time
whole blood takes to coagulate. PTT is often used as a starting
place when investigating the cause of a bleeding or thrombotic
episode. The PTT test is used to determine the effectiveness of
oral anti-coagulation therapy (e.g. Heparin) prescribed to patients
with perturbations in the intrinsic pathway. The above coagulation
tests do not directly measure the effectiveness of drugs (Warfarin,
Coumarin, Heparin etc.) used on blood viscosity (i.e. thinning or
reducing blood viscosity) but measures their second order effect
i.e. blood clotting. The tests conducted in clinics require large
samples of blood (3-5 ml) requiring the addition of
anti-coagulants, with high turn-around times of 1-2 hours. The
currently existing hand-held point of care units for home
coagulation monitoring (COAGUCHEK.RTM., HEMOSENSE.RTM. etc.),
follow the pin-prick blood sampling and strip method as in the
blood-glucose meters, and measure blood coagulation times (PT and
PTT). Though these devices are portable and easy to use, they are
expensive (upwards of $1500/unit and .about.$4/strip) and do not
measure the effect of the anti-coagulation therapy on the actual
physical quantity of blood being affected, i.e. the whole blood
viscosity. A real-time measurement of the physical property of the
complex fluid (here, blood viscosity) can help give a real-time
feedback on the effectiveness and response time of the
treatment/therapy in a clinic allowing for tighter control. Such a
device could in effect be used for measuring standardized PT and
PTT coagulation times for home monitoring as well, thus giving a
complete picture of the therapy induced changes to blood.
[0189] Whole Blood and Plasma Viscosity Measurement
[0190] Red blood cells (RBCs), which compose about 35-40% of the
volume of mammalian blood, are known to be deformable and
contribute to the non-Newtonian nature of blood, i.e., its
viscoelastic behavior (D. R. Gross, N. H. Hwang, "The Rheology of
Blood, Blood Vessels and Associated Tissues," Sitjthoff and
Noordhoff, Rockville, Md., 1981). Whole blood viscosity (WBV) is
observed to have shear-rate dependence due to the deformability of
the RBCs. This has been shown to have a strong correlation to
hypertension (R. B. Devereux, D. B. Case, M. H. Alderman, T. G.
Pickering, S. Chien, J. H. Laragh, "Possible Role of Increased
Blood Viscosity in the Hemodynamics of Systemic Hypertension," Am.
J. Cardiol., vol. 85, pp. 1265-1268, 2000). Also, since the blood
volumes available from patients are small, they must be analyzed
quickly preferably without the addition of anticoagulants. The
currently existing methods for clinical diagnosis and in vitro
study of blood in laboratories, involve the addition of
anti-coagulants thus deviating from the true physiological state of
blood (W. I. Rosenblum, "In vitro measurements of the effects of
anticoagulants on the flow properties of blood: The relationship of
these effects to red cell shrinkage," Blood, vol. 31, no. 2, pp.
234-241, 1968). The plasma viscosity of blood also measured in
clinical laboratories requires the separation of RBCs from blood,
requiring post-processing of blood and delayed response times.
Also, the instruments being bulky pose a practical difficulty in
order to have whole blood and plasma viscosity as a point-of-care
measurement for rapid on-site diagnosis. There is a current need
for low-sample volume (<1 .mu.l), rapid real-time measurement of
rheological properties of whole blood and plasma (viscosity and
coagulation) in vitro or in vivo. Such an instrument, together with
biosensors such as glucose measurement for diabetic patients, could
serve as an invaluable tool for rapid diagnosis and monitoring of
disease and blood function.
[0191] Acoustic wave sensors (piezoelectric crystals and
electro-ceramics) have been used extensively in the measurement of
fluid viscosity (H. Muramatsu, M. Suda, T. Ataka, A. Seki, E.
Tamiya, I. Karube, "Piezoelectric resonator as a chemical and
biochemical sensing device," Sens. Act., vol. 21A, no. 1-3, pp.
362-368, 1990; S. J. Martin, A. J. Ricco, and R. C. Hughes,
"Acoustic wave devices for sensing in liquids," 4th International
Conference on Solid-State Sensors and Actuators (Transducers'87),
Tokyo, Japan, pp. 478-481, June 1987; B. A. Martin, S. W. Wenzel,
and R. M. White, "Viscosity and density sensing with ultrasonic
plate waves," Sens. Act., vol. 22A, no. 1-3, pp. 704-708, 1990).
The physical properties of the fluid can be extracted from the
change in resonant frequency, quality factor and the magnitude of
vibration of a resonator. Thickness-shear mode (TSM) resonators
have been reported to be used in blood rheological characterization
such as viscosity and coagulation measurement. Since the acoustic
shear wave thickness in the blood in contact with the TSM (high
frequency) resonators is much smaller than the size of a RBC, only
plasma viscosity can be measured and the effect of RBCs is not
sensed (H. L. Bandey, R. W. Cernosek, W. E. Lee III, L. E.
Ondrovic, "Blood rheological characterization using the
thickness-shear mode resonator", Biosens. and Bioelec., vol. 19,
pp. 1657-1665, 2004; T.-J. Cheng, H.-C. Chang, T.-M. Lin, "A
piezoelectric quartz crystal sensor for the determination of
coagulation time in plasma and whole blood", Biosens. and Bioelec.,
vol. 13, no. 2, pp. 147-156, 1998). At the opposite extreme are low
frequency resonant devices with high shear wave thickness that can
become bulky, with the size approaching that of commercial tabletop
viscometers. A mid-frequency (30-100 kHz) resonant viscosity sensor
is provided that generates high amplitude motion (>1-micron) via
nonlinear coupling between a longitudinal to transverse motion to
measure fluid viscosity with a self-calibrating immersion depth
sensor. This instrument is likely to be clinically suitable, as a
probe that can be easily inserted in microliters of sample
collected from patients, such as a drop of blood (FIGS. 1, 7). It
has been demonstrated previously (A. Ramkumar, X. Chen A. Lal,
"Silicon Ultrasonic Horn Driven Microprobe Viscometer," IEEE
Ultrasonics Symposium 2006, pp. 1449-1452, 2006) that a
high-amplitude horn-shaped longitudinal mode resonator's quality
factor is sensitive to fluid viscosity into which the probes are
inserted. However, this capability is limited as it is not easy to
measure how much of the probe is inserted into the fluid and, the
shear-rate dependent fluid viscosity cannot be distinguished. Here
fluid viscosity measurement is accomplished by monitoring the
viscous damping of non-linear flexural vibrations of the
microprobes, induced by the .lamda./2 longitudinal resonance of
silicon horn actuators. Real-time coagulation in a rat-blood
droplet (.about.5 .mu.l) as a function of time is measured by
monitoring the resonance frequency of the flexural vibration. This
demonstrates the sensor's applicability as a rapid blood
coagulation sensor alongside measuring whole blood viscosity.
Furthermore, since the viscosity is a strong function of probe
insertion depth, capacitance-based immersion depth sensors have
been integrated into the viscosity sensor for self-calibration of
probe depth, or fluid immersion, of the ultrasonic viscosity
sensor.
[0192] Design and Fabrication of Viscosity Sensor
[0193] The design and fabrication of the microprobe is similar to
the ultrasonic microprobes reported earlier (A. Ramkumar, X. Chen
A. Lal, "Silicon Ultrasonic Horn Driven Microprobe Viscometer,"
IEEE Ultrasonics Symposium 2006, pp. 1449-1452, 2006). In the
embodiment described in the present example, the silicon
microprobes are 4 mm long, 100 .mu.m wide and 140 .mu.m thick
projecting outwards at the tip of the horn (FIG. 1). The
polysilicon strain gauges and immersion depth capacitance sensor
have been integrated to form a multi-sensor microprobe. Dual
multi-sensor microprobes are integrated on the actuator (silicon
horn) to increase the sampling of viscosity measurement but the
measurement can be performed by using a single microprobe.
Polysilicon strain gauges (.about.10 k.OMEGA.) are connected in a
Wheatstone bridge 250 configuration, are integrated at the junction
of the horn and the microprobe to measure the longitudinal and
flexural oscillating motion (FIGS. 1, 2). The silicon horn actuated
at its .lamda./2 longitudinal resonance (109.8 kHz) exerts
longitudinal strain at the tip of the horn, and induces a
non-linear flexural vibration mode (30.4 kHz) in the microprobes.
From the Fast-Fourier Transform (FFT) of the amplified Wheatstone
bridge voltage, the magnitude of vibration corresponding to both
the oscillation frequencies is measured (FIG. 2A). A 2D laser
Doppler vibrometer scan of microprobes indicates a mode with three
displacement nodes with .lamda..about.2.83 mm (FIG. 2B). Immersion
depth calibration is implemented by measuring the capacitance at
discrete steps, which provide distinct motion artifacts as the
liquid with the much higher dielectric constant surrounds the
traces. The distance coding capacitance is formed by platinum
traces (10 .mu.m wide and 0.2 .mu.m thick) co-fabricated on the
microprobes (FIG. 1). FIG. 3 shows the linear dependence of the
capacitance signal with increasing depth of immersion of the
microprobes in saline solution (sensitivity=4.3 nF/mm).
[0194] Results
[0195] Ethylene Glycol Viscosity Measurement
[0196] The flexural vibration amplitude was monitored as the
microprobes were immersed in ethylene glycol at different immersion
depths and was observed at increasing piezoelectric transducer
(PZT) actuation voltages (1-10 V.sub.pp). With increasing immersion
depth of microprobes in deionized water, the damping is observed to
increase with increasing contact with fluid (FIG. 4). FIG. 5 shows
the decrease in vibration amplitude at 8V.sub.pp PZT actuation as
the microprobe is immersed in solutions of varying viscosities. As
a first approximation, the viscous fluid damping of the microprobe
is modeled by assuming the immersed part of the vibrating probe to
be an oscillating sphere immersed in a liquid (C. Riesch, E. K.
Reichel, F. Keplinger and B. Jakoby, Journal of Sensors, Article ID
697062, 2008), and the damping model was shown to agree reasonably
with the data in FIG. 5.
[0197] The parameters in the model are dependent of the microprobe
mass and geometry, and the immersion depth. Therefore, a wide range
of viscosities can be accurately measured by monitoring the
amplitude of the vibration in the viscosity sensor at different
immersion depths (in this embodiment, both capabilities are
integrated on the device).
[0198] The performance of the viscosity sensor can be assessed by
calibrating it using a standard, such as ethylene glycol, with a
known viscosity at a given temperature. The quality of the
analytical damping model's fit to data obtained from the viscosity
sensor can be used to assess its performance. Fluid viscosity can
also be determined by monitoring the change in frequency (f.sub.0)
and quality factor (f.sub.0/.DELTA.f) of the flexural and
longitudinal vibrations. This can permit higher sensitivity when
measuring fluid viscosity.
[0199] Measurement of Rat Whole Blood Coagulation
[0200] The viscosity sensor was used to measure rat whole blood
coagulation. Whole blood was obtained from a Sprague-Dawley rat (44
weeks old) using exsanguination. No anti-coagulants were added. The
flexural vibration frequency was monitored as the microprobe of the
viscosity sensor was immersed in the blood at a depth of .about.1
mm. Initially, the red blood cells (RBCs) coated the microprobe,
lying within the shear thickness and thus loading the vibration by
reducing its frequency (Phase I, FIG. 6). The blood coagulation
cycle ramped up by the formation of blood clots (coagulated RBCs)
with sizes bigger than the shear thickness, and thus reducing the
loading on the vibration frequency (Phase II, FIG. 6). Finally, the
blood coagulation cycle ended at .about.430 seconds (Phase III,
FIG. 6), which was close to that known in the art (T.-J. Cheng,
H.-C. Chang, T.-M. Lin, "A piezoelectric quartz crystal sensor for
the determination of coagulation time in plasma and whole blood",
Biosensors and Bioelectronics, vol. 13, no. 2, pp. 147-156, 1998).
The residual damping observed at the end of the coagulation cycle
was mostly due to the plasma; few RBCs and the fibrous network of
clots formed. The microprobe was cleaned afterwards by immersing in
diluted bleach solution and actuating the PZT at 20V.sub.pp, and
the frequency and amplitude of vibration returned to the baseline
values.
[0201] These experiments demonstrate that fluid viscosity can be
measured by monitoring the flexural vibration of the immersed
microprobes, with precise control of the depth of immersion. Since
the acoustic shear thickness of the flexural vibration is close to
the size of a RBC (C. Riesch, E. K. Reichel, F. Keplinger, B.
Jakoby, "Characterizing Vibrating Cantilevers for Liquid Viscosity
and Density Sensing," J. Sensors, Article ID 697062, pp. 1-9,
2008), the viscous damping was influenced by the RBCs, thus
enabling the measurement of whole blood viscosity. Also, by varying
the PZT actuation voltage the microprobe flexural vibration rate
can be varied, and hence shear rate dependence of whole blood
viscosity can be measured. In addition, the damping in the silicon
horn's longitudinal vibration (109.8 kHz) having a shear thickness
much smaller than the size of a RBC will allow for measurement of
blood plasma viscosity (A. Ramkumar, X. Chen A. Lal, "Silicon
Ultrasonic Horn Driven Microprobe Viscometer," IEEE Ultrasonics
Symposium 2006, pp. 1449-1452, 2006). Since the shear (non-linear
vibration) and bulk (silicon horn's .lamda./2 longitudinal
resonance) damping can be observed at two different frequencies,
one can simultaneously measure the plasma (bulk) and whole
(shear-rate dependent) blood viscosity in low-sample volumes
(.about.1-3 .mu.l). In addition, the natural and drug-induced blood
coagulation cascade can be precisely monitored as a function of
time. Circuitry can be designed using methods known in the art to
monitor the change in frequency f.sub.0 and quality factor
(f.sub.0/.DELTA.f) of the flexural and longitudinal vibrations of
the microprobes when immersed in a fluid. This enables higher
sensitivity when measuring fluid viscosity.
Example 2
Assessment of Flow Measurement Ability of Piezoresistive
Microprobes
[0202] This example describes an experiment that tested the
flow-sensing properties of an embodiment of the device. This
embodiment of the device comprised piezoresistive microprobes 20
that were positioned upside down and perpendicular into a narrow
channel 300 that was approximately 8 mm width and approximately 5
cm in length. The probe tips extended approximately 2 mm inside the
channel. FIG. 9a shows a photograph of this embodiment of the
device. FIG. 9b shows the tips of the piezoresistive probes inside
the tunnel (channel) 300.
[0203] Water was conducted through a narrow (.about.1 mm in
diameter) tubule 400 inside the channel. The other side of tubule
was connected to a 30 ml syringe while sitting in a syringe driver
600. The driver pushed the syringe to produce a 30 ml/min flow of
water. The probes were connected to an interface to monitor the
induced voltage in the Wheatstone bridge 250 integrated into the
bodies of the microprobes. Such an interface is commonly known in
the art. A 5 kHz signal was used in this interface for data
sampling for period of 20 sec.
[0204] When the syringe driver started to pump water into channel,
it took some time for the system to become stable. Voltage changes
were parabolic in this regime. In addition, the measured voltage
was constant or linear depending on the water depth in channel. If
the water filled the channel gradually during the experiment, a
linear regime in measured voltage was observed. If the channel
saturated from water and became almost constant in depth, the
voltage became constant during the measurements.
[0205] FIG. 10 shows the measured voltage versus time in probes
when the channel was being filled. The syringe driver was turned on
and sampling started. It took about 3.5 seconds for the system to
stabilize and then a linear regime started. In this experiment, the
slope of the linear regime was 5.9 mv/sec.
[0206] Another measurement was then done after a 5-second pause
after switching the syringe driver on to monitor the stable
condition of the system. FIG. 11 plots voltage versus time in this
experiment. In the case of constant channel depth, the same method
can be used to measure probe voltage.
[0207] As shown in FIG. 12, it takes around 3.5 seconds for system
to become stable, then voltage become constant; then voltage
immediately drops to its initial value (-1.88.sup.V). The plot is
not uniformly constant; vibration noise in the environment, e.g.,
syringe driver is a potential noise source. Second, due to small
channel length, water hits the holder wall and bounce back
thereafter and this could cause a voltage drop. Finally, any
turbulence in channel could cause a drop at output voltage. The
average voltage during the experiment was -1.835.sup.V and voltage
drop owing to fluid flow (30 ml/min) was 45 mV (FIG. 12).
[0208] By making a longer channel to decrease settling time of
system, the performance of the system can be improved. Additional
measurements in different flow values can be used to determine a
suitable range of flow measurement for fabricated piezo-resonators.
Performing measurements with the device positioned on a noise
cancellation table can be used to obtain more constant
voltages.
[0209] The present invention is not to be limited in scope by the
specific embodiments described herein. Indeed, various
modifications and variations of the invention in addition to those
described herein will become apparent to those skilled in the art
from the foregoing description and can be made to the invention
without departing from the spirit and scope of the invention. Thus,
it is intended that the present invention cover the modifications
and variations of this invention provided they come within the
scope of the appended claims and their equivalents.
[0210] All references cited herein are incorporated herein by
reference in their entirety and for all purposes to the same extent
as if each individual publication, patent or patent application was
specifically and individually indicated to be incorporated by
reference in its entirety for all purposes.
[0211] The citation of any publication is for its disclosure prior
to the filing date and should not be construed as an admission that
the present invention is not entitled to antedate such publication
by virtue of prior invention.
* * * * *